A recent publication indicates that the diagnostic performance of first-pass cardiac perfusion MRI is better at 3 T than at 1.5 T, for the identification of both single and multiple vessel coronary disease (see, e.g., Cheng et al., Cardiovascular magnetic resonance perfusion imaging at 3-tesla for the detection of coronary artery disease: a comparison with 1.5-teslam, Journal of the American College of Cardiology; 49(25), 2440-2449 (2007)). However, transmit radio-frequency (RF) field (B1+) variations and static magnetic field (B0) variations can be comparatively higher at 3 T than at 1.5 T (see, e.g., (see, e.g., Greenman R L et al., Double inversion black-blood fast spin-echo imaging of the human heart: a comparison between 1.5T and 3.0T, Journal of Magnetic Resonance Imaging 2003; 17(6):648-655.; Noeske R, Seifert F et al., Human cardiac imaging at 3 T using phased array coils. Magnetic Resonance in Medicine 2000; 44(6):978-982; Singerman R W et al., Simulation of B1 field distribution and intrinsic signal-to-noise in cardiac MRI as a function of static magnetic field, Journal of Magnetic Resonance 1997; 125(1):72-83; Schar M et al., Cardiac SSFP imaging at 3 Tesla, Magnetic Resonance in Medicine 2004; 51(4):799-806; Sung K, Nayak K S. Measurement and characterization of RF nonuniformity over the heart at 3T using body coil transmission, Journal of Magnetic Resonance Imaging 2008; 27(3):643-648). Specifically, the peak-to-peak variation of B0 within the heart at 3 T has been reported to be on the order of 130-260 Hz (see, e.g., Noeske R et al., supra.), and B1+ variation within the heart at 3 T has been indicated to be on the order of 50-80% (see, e.g., Singerman R W et al., supra.; Sung K et al., Measurement and characterization of RF nonuniformity over the heart at 3T using body coil transmission, Journal of Magnetic Resonance Imaging 2008; 27(3):643-648). These challenging factors make it more difficult to perform uniform T1-weighting at 3 T using a conventional non-selective saturation (e.g., 90°) pulse, which is known to be sensitive to both B0 and B1+ inhomogeneities (see, e.g., Ernst RR et al., Bodenhausen G, Wokaun A. Off-resonance effects due to finite pulse amplitude. In: Principles of nuclear magnetic resonance in one and two dimensions. New York: Oxford University Press Inc.; 1987. p 119-124; Kim D, Cernicanu A et al., B(0) and B(1)-insensitive uniform T(1)-weighting for quantitative, first-pass myocardial perfusion magnetic resonance imaging, Magnetic Resonance in Medicine 2005; 54(6):1423-1429; Kim D, Gonen O et al., Comparison of the effectiveness of saturation pulses in the heart at 3T, Magnetic Resonance in Medicine 2008; 59(1):209-215).
Non-uniform T1-weighting can be disadvantageous for both the interpretation and quantitative analysis of first-pass cardiac perfusion MR images. Given the limitations of commercially available automated B0 shimming and RF calibration procedures for the heart, it can be important to design robust saturation pulses that can be insensitive to clinically relevant B0 and B1+ variations within the heart at 3 T. Previously proposed adiabatic B1-insensitive rotation (BIR-4) (see, e.g., Kim D, Gonen O et al., supra.; Staewen R S et al., 3-D FLASH imaging using a single surface coil and a new adiabatic pulse, BIR-4, Investigative Radiology 1990; 25(5):559-567) pulse, RF pulse train of three non-selective 90° pulses (see, e.g., Kim D, Gonen O et al., supra.; Oesingmann N et al., Improved saturation RF pulse design for myocardial first-pass perfusion at 3T, Journal of Cardiovascular Magnetic Resonance 2004; 6(1):373-374) and RF pulse train of three non-selective pulses with different flip angles (e.g., 96°, 228°, 141°) (see, e.g., Sung K et al., Design and use of tailored hard-pulse trains for uniform saturation of myocardium at 3 Tesla, Magnetic Resonance in Medicine 2008; 60:997-1002) can be used to improve the efficacy compared with a conventional non-selective 90° pulse, but at the expense of relatively longer pulse duration and higher specific absorption rate (SAR). For convenience, the pulse train of three non-selective 90° pulses can be referred to as the standard pulse train, and the pulse train of three non-selective pulses with different flip angles will be referred to as the tailored pulse train. The longer pulse duration (10-12 ms) can be inconsequential because a typical perfusion image acquisition time is on the order of 120-150 ms.
A higher SAR can likely be a significant limitation for a BIR-4 pulse (see, e.g., Kim D, Gonen O et al., supra.) because it can limit the nominal B1+ of the pulse and also reduces the slice coverage. An advantage of both standard and tailored pulse trains can be that their RF deposition levels can be considerably lower than that of a BIR-4 pulse (see, e.g., Kim D, Gonen O et al., supra.; Sung K et al., Design and use of tailored hard-pulse trains for uniform saturation of myocardium at 3 Tesla, Magnetic Resonance in Medicine 2008; 60:997-1002). A standard pulse train can be used to perform uniform saturation of magnetization within the left ventricle (LV), but not within the right ventricle (RV), where the B1+ variation is comparatively higher (see, e.g., Greenman R L et al., supra.); Kim D, Gonen O et al., supra.; Sung K et al., Design and use of tailored hard-pulse trains for uniform saturation of myocardium at 3 Tesla, Magnetic Resonance in Medicine 2008; 60:997-1002). As an extension of the WET (water suppression enhanced through T1 effects) pulse (see, e.g., Ogg R J et al., WET, a T1- and B1-insensitive water-suppression method for in vivo localized 1H NMR spectroscopy, Journal of Magnetic Resonance Series B1994; 104(1):1-10), a particular publication (see, e.g., Sung K et al., Design and use of tailored hard-pulse trains for uniform saturation of myocardium at 3 Tesla, Magnetic Resonance in Medicine 2008; 60:997-1002) describes a tailored pulse train to numerically optimize the three flip angle values in order to minimize the residual longitudinal magnetization (MZR), which can be calculated as longitudinal magnetization (Mz) divided by the equilibrium magnetization (M0). The tailored pulse train design was likely based on prior measurements of B0 ranging from ±140 Hz and B1+ scale ranging from 0.4-1.2 (i.e., B1+ scale=1.0 represents true B1+). However, this can be at the expense of producing medium level of non-uniform residual magnetization within this range (see FIG. 3 in Ref. 12; see FIGS. 2-3) and slightly higher SAR than that of a standard pulse train (see, e.g., Sung K et al., Design and use of tailored hard-pulse trains for uniform saturation of myocardium at 3 Tesla, Magnetic Resonance in Medicine 2008; 60:997-1002).
Despite such improvements, these pulses described in such publications are likely incapable of performing complete saturation of magnetization within the whole heart at 3 T while remaining within clinically acceptable SAR limits. One of the objects of the present disclosure is to provide a hybrid adiabatic-rectangular pulse train that can achieve both of the aforementioned objectives at 3 T and evaluate its efficacy and predicted SAR against those of the standard pulse train, tailored pulse train, and BIR-4 pulses.